The present invention relates generally to the field of medical diagnostic imaging. More particularly, the present invention relates to the magnetic resonance imaging and to the reduction of noise within magnetic resonance imaging systems incorporating ultra-short cylindrical magnets.
Magnetic resonance imaging (MRI) systems have become ubiquitous in the field of medical diagnostics. Over the two past decades, improved techniques for MRI examinations have been developed that now permit very high-quality images to be produced in a relatively short time. As a result, diagnostic images with varying degrees of resolution are available to the radiologist that can be adapted to particular diagnostic applications.
In general, MRI examinations are based on the interactions among a primary magnetic field, a radiofrequency (rf) magnetic field and time varying magnetic gradient fields with nuclear spins within the subject of interest. Specific nuclear components, such as hydrogen nuclei in water molecules, have characteristic behaviors in response to external magnetic fields. The precession of spins of such nuclear components can be influenced by manipulation of the fields to obtain rf signals that can be detected, processed, and used to reconstruct a useful image.
The magnetic fields used to produce images in MRI systems include a highly uniform, static magnetic field that is produced by a primary magnet. A series of gradient fields are produced by a set of three gradient coils disposed around the subject. The gradient fields encode positions of individual volume elements or voxels in three dimensions. A radiofrequency coil is employed to produce an rf magnetic field. This rf magnetic field perturbs the spin system from its equilibrium direction, causing the spins to precess around the axis of their equilibrium magnetization. During this precession, radiofrequency fields are emitted by the spins and detected by either the same transmitting rf coil, typically a birdcage resonator, or by a separate receive-only coil. These signals are amplified, filtered, and digitized. The digitized signals are then processed using one of several possible reconstruction algorithms to reconstruct a useful image.
Many specific techniques have been developed to acquire MR images for a variety of applications. One major difference among these techniques is in the way gradient pulses and rf pulses are used to manipulate the spin systems to yield different image contrasts, signal-to-noise ratios, and resolutions. Graphically, such techniques are illustrated as “pulse sequences” in which the pulses are represented along with temporal relationships among them. In recent years, pulse sequences have been developed which permit extremely rapid acquisition of a large amount of raw data. Such pulse sequences permit significant reduction in the time required to perform the examinations. Time reductions are particularly important for acquiring high-resolution images, as well as for suppressing motion effects and reducing the discomfort of patients in the examination process.
While field interactions are fundamental to the encoding of data acquired in MRI systems, certain field interactions are undesirable, or may lead to degradation of the image data. For example, when the appropriate pulses are applied to an rf coil during an examination sequence, rf energy from the rf coil can penetrate the gradient coil structure where it is dissipated by lossy eddy currents induced in the gradient coil structure. To maintain a high efficiency of the rf coil, an rf shield is typically positioned between the rf coil and the gradient coil set so as to prevent or reduce penetration of the rf magnetic field into all of the gradient coils. The design of the rf shield is such that minimal eddy currents are generated by switching of the gradient fields, rendering the rf shield substantially transparent to the gradient fields. At the same time, the rf frequencies are much higher than characteristic eddy current decay rates in the shield, hence the shield functions like an impenetrable barrier to rf fields. However, the proximity of an rf shield to the rf coil conductors, particularly in the case of a whole body rf transmit coil, may significantly affect the overall power efficiency and the signal-to-noise ratio of the rf coil. Therefore, in general, it is desirable to place the rf shield as far as possible from the rf coil.
To address these concerns, the rf shield may be placed between the gradient coils such that the Z-axis gradient coil, typically an antisymmetric solenoid-type coil of varying pitch, is positioned within the shield, that is, between the shield and the rf transmit coil. This configuration is possible because the mode of the rf coil that is typically used in MRI has little or no net magnetic flux in the Z-axis direction, resulting in minimal coupling between the rf coil and the Z-axis gradient coil. In this configuration, the radiofrequency field is essentially undisturbed by the presence of the Z-axis gradient coil on the interior of the shield surface, allowing the rf shield to be moved significantly away from the transmit coil and thereby providing a significant reduction in noise and an increase in efficiency.
This configuration, however, may be unacceptable in systems using ultra-short cylindrical magnets where the spacing between the windings of the Z-axis gradient coil is drastically reduced. At this reduced spacing, a number of factors appear to prevent proper decoupling of the high-density z-gradient coil from the rf coil. First, in such systems the distance between the rf coil conductors and the Z-axis gradient conductors may be as little as 10 mm, resulting in frequent misalignment, either radially or in the direction of the Z-axis, which leads to coupling between the rf coil and the Z-axis gradient coil. Second, coupling occurs between the end rings of the birdcage resonator and the Z-axis gradient windings as they pass over the end rings. Finally, the quality factor, or Q, associated with the system depends upon the angle between the conductive “rungs” of the birdcage resonator and the connecting wire which connects the 2 antisymmetric halves of the Z-axis gradient coil. In particular, Q is reduced each time the connecting wire passes above one of the birdcage rungs.
There is a need, therefore, for an improved technique for winding a Z-axis gradient. To address the drawbacks in hereto foreknown systems, there is a particular need for a technique which minimizes interactions between the z-gradient and the birdcage rf body coil in systems employing ultra-short cylindrical magnets and a rf shield outside of the Z-axis gradient coil.